1. Field of the Invention
The present invention relates to an apparatus for nuclear spin tomography that uses superconducting magnetic coils to generate a homogenous magnetic base field. Normal conducting coils within the interior of the superconducting base field coils develop magnetic field gradients. At least one cooled radiation shield of electrically and thermally conducting material is arranged between the normal conducting gradient coils and the superconducting base field coils.
2. Description of Related Art
Image producing diagnostic methods have been developed in the field of medical technology to measure integral resonance signals of nuclei of a given chemical element in a body or part of a body such as a human body. The spatial spin density or relaxation time distribution obtained in this manner can be used to construct an image similar to an x-ray tomogram in computer tomography. Similar methods are generally known under the designation "nuclear spin tomography" or Nuclear Magnetic Resonance (NMR) tomography, also called "Magnetic Resonance Imaging (MRI) or Magnetic Resonance Spectroscopy (MRS).
Nuclear spin tomography requires a magnetic field that is produced by a base field magnet. A body or the part of a body to be examined is placed along an axis that generally coincides with the orientation axis of the magnetic base field. The base field must be sufficiently homogeneous in a suitable imaging or examination zone. The magnetic induction in this zone can amount to several Tesla. Such high magnetic induction, however, can be generated only with superconducting magnet coils which must be accommodated in a suitable cryo system comprising at least one cooled radiation shield that limits heat input from the room temperature that reaches the superconducting base-field coils. The design of the radiation shield is generally governed by thermal considerations and made of highly electrically conductive material. Stationary or pulsed gradient fields are superimposed on the magnetic base field. The superimposed field is formed by nonsuperconducting coils that are arranged within the interior of an area defined by the cryo system of the base field coils and generally operate at approximately room temperature. A separate antenna is further require to excite the individual atom nuclei in the body or the part of a body to perform a precession motion. Thus, a high frequency a-c magnetic field can be generated for a short time. Optionally, this antenna can also receive the high frequency signals generated by the excited atomic nuclei.
According to the European Patent Application 0,144,171, the radiation shield of a superconducting base field coil system or its cryo system is made of a electrically highly conducting material such as aluminum. The nonsuperconducting gradient coils, however, generates eddy currents in the radiation shield that produce reaction gradient fields in the useful volume of the examination zone. These eddy currents decay with a time constant .tau. which is generally between 0.05 and 0.3 seconds due to the finite conductivity of the radiation shield. This leads to a pulse distortion of the gradient field, or a frequency response curve of the sensitivity of the coil system defined as the gradient field strength per operating current which increases below a frequency limit f.sub.g toward low frequencies f&lt;f.sub.g =(2.pi..tau.).sup.-1. The eddy currents thus reduce resolution in nuclear spin tomography unless countermeasures are taken. It is generally not feasible to make the radiation shields non-shielding by using, for example, poorly conductive material or slots. The gradient fields can penetrate these shields and generate heat in the helium chilled region of the superconducting base field coils. This heating is particularly disturbing at their very low operating temperature.
Primarily two countermeasures are known in the art. The first method has the magnetic field of the eddy currents induced in the radiation shield at about 10-30% of the direct gradient field. The magnetic field of the eddy currents forms its own gradient field which is opposed to the original field and therefore attenuates it. This effect can be taken into consideration in designing the gradient coils to produce a combined gradient field of high spatial quality. (See, "J. Phys. E", Vol. 19, 1986, pages 876 to 879). The frequency behavior of the eddy currents can be compensated for by using a frequency correction of the current pulses that feed the gradient coils. In general, however, such a correction is unique to each magnet system since the conductivity of the laminations used as the radiation shield is scattered at low temperatures. In addition, it cannot be assured that the spatial distribution of the eddy currents remains unchanged during the decay. The separability of the space and time behavior of the gradient field with respect to the radiation shield, assumed with this countermeasure, does not always exist.
A second countermeasure it is known, for example, from "J. Phys. D", Vol. 19, 1986, pages L129 to L131. This countermeasures uses an additional system of gradient coils that are positioned between the primary gradient coil system and the radiation shield and positioned as close as possible to the shield. It is possible to produce an improved frequency curve which is at least largely smooth. The crowded space conditions, however, in the interior of the solenoid-like superconducting base field coil lead to difficulties. Each negatively excited additional gradient coil is located very close to the corresponding primary coil since its radii must be only about 20% larger in radius. As a consequence, the gradient field action is largely cancelled. Therefore, distinctly higher supply voltages and currents are necessary than with the countermeasures explained above because a magnet coil system designed according to this countermeasure starts eddy currents in the cold, highly conductive radiation shield. These eddy currents cost less Joule power than the currents in the additional gradient coils that are at approximately room temperature. The inductive reactive power is smaller with the first countermeasure because the radiation shield is located radially further outward than the additional gradient coils used in the second countermeasure. Present and future pulse sequences, however, require even stronger gradient fields, especially for fast image generation. The electronic power circuitry for feeding the gradient coils becomes a correspondingly more expensive part of the overall system, especially for using the second countermeasure.